Image guidance for radiation therapy is an active area of investigation and technology development. Current radiotherapy practice utilizes highly conformal radiation portals that are directed at a precisely defined target region. This target region consists of the Gross Tumour Volume (GTV), the Clinical Target Volume (CTV) and the Planning Target Volume (PTV). The GTV and CTV consist of gross tumour disease and the subclinical microscopic extension of the gross disease. During radiation treatments, these volumes must be irradiated at a sufficient dose in order to give an appropriate treatment to the patient. Because of the uncertainty in identifying this volume at the time of treatment, and due to unavoidable patient and tumour motion, an enlarged, the PTV is typically irradiated.
Because a volume that is larger than the biological extent of the disease is typically irradiated, there is an increased risk of normal tissue complications due to the unnecessary irradiation of healthy tissue. Thus, it is desirable to conform the radiation beam to the GTV and CTV only, and to provide an imaging method to assist in the placement of the radiation beam on this volume at the time of treatment. This technique is known as Image Guided Radiation Therapy (IGRT).
Commercially available techniques that are available for IGRT typically use x-ray or ultrasound imaging technology to produce planar x-ray, computed tomography, or 3D ultrasound images. Furthermore, fiducial markers can be used in conjunction with these imaging techniques to improve contrast. However, fiducial markers must be placed using an invasive technique, and are thus less desirable. IGRT techniques based or x-rays or ultrasound are not ideally suited to IGRT: x-rays suffer from low soft tissue contrast and are not ideally suited to imaging tumours; ultrasound cannot be utilized in all locations of the body. Further, x-ray based techniques use ionizing radiation and thus deposit supplemental dose to the patient. Finally, x-ray and ultrasound based IGRT techniques are difficult to integrate into a linear accelerator such that they can provide images in any imaging plane in real time at the same moment as the treatment occurs.
In order to overcome these difficulties, it has been proposed to integrate radiotherapy systems with a Magnetic Resonance Imaging (MRI) device. As is well known, MRI offers excellent imaging of soft tissues, and can image in any plane in real time.
An MRI functions by providing a homogeneous, and strong magnetic field that aligns the nuclear magnetic moments of target nuclei; hydrogen nuclei (protons) are the most common imaging target in MRI. In the presence of the magnetic field, the magnetic moments of the nuclei align with the homogeneous magnetic field and oscillate at a frequency determined by the field strength; this frequency is known as the Larmor frequency. This alignment can be perturbed using a radiofrequency (RF) pulse, such that the magnetization flips from the direction of the magnetic field (B0 field) to a perpendicular direction, and thus exhibits transverse magnetization. When the nuclei reverts back to its original state, the transverse magnetic moment decays to zero, while the longitudinal magnetic moment increases to its original value. Different soft tissues exhibit different transverse and longitudinal relaxation times. A specific magnetic field strength is applied to a small sample of tissue utilizing gradient magnetic coils, and images of these soft tissues can be formed by generating a specific sequence of perturbing RF pulses and analyzing the signals that are emitted by the nuclei as they return to their original magnetization state after being perturbed by the first RF pulse.
A medical linear accelerator functions by using a cylindrical waveguide that is excited in a TM010 mode such that the electric field lies upon the central axis of the waveguide. The phase velocity of the structure is controlled by introducing septa into the waveguide which form cavities. The septa have small holes at their centre to allow passage of an electron beam. Septa have the further advantage that they intensify the electric field at the center of the waveguide such that field gradients in the MeV/m range are available for RF input power that is in the MW range. Electrons are introduced into one end of the accelerating structure, and are then accelerated to MeV energies by the central electric field of the accelerating waveguide. These electrons are aimed at a high atomic number target, and the electronic energy is converted in high energy x-rays by the bremsstrahlung process. The waveguide is typically mounted on a C-arm gantry such the central axis of the waveguide is parallel to the ground. This waveguide rotates around a patient, which lies at the central axis of rotation. The medical accelerator utilizes a system employing a 270° bending magnet such that the radiation beam generated by the waveguide is focused at a point on the central axis of rotation known as the isocentre.
As is known, there are several significant technological challenges associated with the integration of a linear accelerator with an MRI device.
For example, if the linear accelerator is physically close to the MRI, the large magnetic field of the MRI magnet can affect the acceleration of electrons in the accelerating waveguide since electrons axe charged panicles, and are thus influenced by the Lorentz force, F=q(v×B), where v×B is the cross product between the electron velocity v, and the magnetic flux density B. If the direction of the electron motion is perpendicular to the magnetic field direction, the deflection of the electron's path will be a maximum, and it will very often result in electrons tending to collide with the side wall of the linear accelerator, which will stop the particle accelerating process.
A further challenge is due to the pulsed power nature of the linear accelerator. In order to supply sufficient RF power (on the order of Mega-Watts MWs) to the accelerating waveguide, medical linear accelerators operate in a pulsed power mode where high voltage is converted to pulsed power using a pulse forming network (PFN). The process of generating high voltage pulses involves sudden starting and stopping of large currents in the modulation process, and these in turn can give rise to radiofrequency emissions whose spectrum can overlap the Larmor frequency of the hydrogen nuclei within the imaging subject. This would thus interfere with the signals emitted by these nuclei as they relax, and would thus deteriorate the image forming process of the MRI.
A further challenge that exists when integrating an MRI with a medical linear accelerator involves the orientation of the MRI magnet with respect to the accelerating waveguide such that the waveguide can be directed at the patient without obstruction from the magnet.
A further challenge relates to the dose deposition pattern obtained when a patient is exposed to high energy x-ray used in radiotherapy in the presence of a strong magnetic field used far imaging by MRI. The dose deposited by the MRI is due to electrons scattered by the incoming photons by the photoelectric, Compton, or pair production processes. These electrons are charged particles, and are also subject to the Lorentz force. If the direction of the magnetic flux density is perpendicular to the incident direction of the x-ray beam, this produces perturbations to the dose deposition pattern that are significant, and increase in magnitude as the magnetic flux density increases.
U.S. Pat. No. 6,365,798 to Green discloses a method of mounting an open, bi-planar magnet on a conventional C-arm medical linear accelerator. The design of the linear accelerator is not changed from that built by the patent assignee (Varian), and much of the patent describes methods of retrofitting an MRI magnet to an existing design of linear accelerator. Several configurations of an MRI magnet are described. For example, the magnet can be mounted independently of the C-arm accelerator, and thus remain stationary. In this configuration the magnet has a wide enough opening to allow irradiation from several angles. Alternatively, the MRI magnet is mounted on the C-arm gantry, and rotates with the gantry to provide rotational therapy. Green relies on the MRI magnet being small enough to be able to be added to an existing medical accelerator manufactured by Varian Medical Systems.
Further, according to Green the MRI magnet is positioned so that the radiation beam itself is parallel or perpendicular to the direction of the main magnetic field. Several magnet orientations are described, and include coils with a central opening (for passage of the patient or the radiation beam) or no central opening.
To avoid interference between electron acceleration or the linear accelerator's 270° bending magnet and the MRI, Green suggests a low magnetic field that is only just sufficient to provide the lowest quality image to align a beam with a specified region of tissue. As well, active shielding methods to reduce the magnetic field from the MRI magnet at the accelerating waveguide of the medical linear accelerator are employed. Green does not contemplate a solution to RF interference problems described above.
The Green document also suggests a mechanism whereby an x-ray beam would cause spectral changes in the NMR spectra of the tissues being irradiated, and further describes a method that would image the region of tissue being irradiated based on the suggested NMR spectral changes.
PCT Patent Application Publication No. WO 2004/024235 to Lagendijk discloses a cylindrical, solenoid shaped MRI magnet that is combined with a linear accelerator that is mounted perpendicular to this magnet (at its mid-point) and points to the central axis of the magnet. The patient lies on the central axis of the cylindrical magnet, with the magnetic field in the cranial-caudal direction. The radiation beam is perpendicular to the direction of magnetic field. The magnet is designed with active shielding such that the magnetic field where the accelerating waveguide is located is reduced to a low value. The radiation beam must penetrate the solenoidal magnet to reach the patient located inside the solenoid, and is thus attenuated by the magnet. Filters are described to compensate for the effects of the solenoidal magnet on the quality of the x-ray beam. As well an embodiment whereby the solenoidal magnet is split such that an unattenuated x-ray beam reaches the patient is also described. No solutions to the RF interference, or perturbation of the dose distribution by the B0 field are described.
U.S. Pat. No. 6,862,469 to Bucholz et al, discloses a method to combine a proton beam with an MRI system. This invention is indirectly related to the current disclosure since it relates to proton therapy, and does not discuss methods of bringing a medical linear accelerator close to an MRI magnet. This disclosure describes a photon beam that impinges through an aperture of an MRI magnet in the same direction as the B0 field, and is thus not deflected by the B0 field since the vector product v×B is 0 in this case. A limitation of this disclosure is the small aperture size in the magnet.
Specific discussion about interference between the manufacture of the proton beam MRI operation is not discussed. Typically, in proton irradiators, the proton beam is accelerated to the desired energy far from the patient. It is thus implied that the proton acceleration process does not produce magnetic interference with the MRI.
A significant part of the Bucholz et al. disclosure relates to feedback methods whereby the MRI imaging information is used to position the proton beam at the suitable position on the patient.
Bucholz et al. describe a system where the patient is rotated is for rotational therapy; however gantry rotation of the proton beam is briefly mentioned. For a rotating gantry, Bucholz describes a stationary MRI magnet where beam access through the magnet gap is proposed.
Bucholz et al. further briefly mention other magnetic and RF interference, and suggests that shielding methods ear be used to remove these, if needed.
PCT Patent Application Publication No. WO 2006/136865 to Kruip et al. discloses a MRI system that can be combined with proton therapy. A sophisticated magnet design is described that allows the proton beam to be in the same direction as the B0 field of the MRI, but with a large opening that would allow for translation of the proton beam. The magnet design is complicated, and involves non-circular and complex coils. As in Bucholz et al., the proton source is far from the magnet, and the two devices are assumed not to interfere with each other magnetically. Furthermore, no discussion of rotation therapy is described.
U.S. Patent Application Publication No. 2005/0197564 to Dempsey discloses a method of delivering radiotherapy using a radionuclide as the source of ionizing radiation in combination with an open solenoid MRI. The patient is place in the bore of the MRI magnet such that the magnetic field is parallel to cranial-caudal direction. The radionuclide used is 60Co and it is placed such the patient is irradiated through the opening of the MRI solenoid, and so the magnetic field is perpendicular to the direction of the x-ray beam. 60Co is radioactive, and emits photons with a mean energy of 1.25 MeV.
In the Dempsey design, no accelerating waveguide is used and so the problems of an electron beam deflection in the accelerating waveguide by the B0 field of the MRI are not encountered. RF interference between a medical linear accelerator and an MRI are also avoided since 60Co does not use a PFN. However, this method introduces a new problem in that 60C is ferromagnetic, and will thus introduce inhomogeneities in the B0 field of the MRI. When the 60Co source is rotated, these inhomogeneities will degrade the MRI image quality, which necessitates the use of novel techniques to recover the image quality. As well, 60Co has a finite dose rate for a given source activity, which reduces in time due to the half life of 60Co. This dose rate is generally lower than that of a medical linear accelerator and is thus undesirable. As well, 60Co source size is large enough such that the focal point of the radiation source is larger than that of a medical linear accelerator. This reduces the x-ray beam quality of the 60Co source as compared to that of the medical linear accelerator.
Dempsey describes in some detail the perturbation effects on the dose distribution in the patient due to the magnetic field. He suggests that at 1.5 T, these perturbations are significant for perpendicular irradiation, but are considerably reduced, if not eliminated, when a low field MRI, such as 0.3 T, is used in perpendicular irradiation.
Dempsey also describes the use of alternate irradiation sources such as protons or neutrons.
PCT Patent Application Publication No WO 2007/045076 to Fallone et al., assigned to the assignee of the present application, and the contents of which are incorporated herein by reference, describes a medical linear accelerator that is combined with a bi-planar permanent magnet suitable for MRI.
While the documents described above provide various advancements, there are technological problems that are yet to be resolved. For example, several of the configurations described above propose a reliance on the use of magnetic shielding, or shimming. Such shielding is strategically placed on or around the system to mitigate the magnetic effect of the MRI on the linear accelerator as or to compensate for the effect of the ferromagnetic 60Co on the MRI. As a result of this reliance on magnetic shielding/shimming, these systems tend to be designed so as to provide as large a distance as possible between the MRI and linear accelerator or 60Co source. Unfortunately, increasing the distance between the MRI and linear accelerator lowers the photon dose rate seen at the MRI isocentre. As a result, the treatment time for providing necessary dosage is prolonged. Another significant drawback to the larger distance required between the linear accelerator and the isocentre is the according increase in the physical size of the combined MRI-linear accelerator. Such increases in size lead to difficulties in ensuring integrated MRI-linac devices can be installed in standard-size radiation therapy suites. As would be understood, a configuration that relies far less or not at all on magnetic shielding/shimming for reducing magnetic interference would pose fewer restrictions on the relative placement of the linac and the MRI. As a result, such a configuration would enable size reduction and dose rate increases without undesirable magnetic interference.
A second difficulty that is common to the above is that these produce dose distributions in the patient that are perturbed from the case where there is no magnetic field. This perturbation is due to the Lorentz force on the scattered electrons that originate when photons interact with the biological material of the patient. One of more of the above proposals use a device arrangement where the photon beam is perpendicular to the B0 field of the MRI imager, and so in this case the Lorentz force on the scattered electrons is greatest. A device where the B0 field was parallel to the direction of the photon beam would produce scattered electrons of which a great majority have a small angle of travel with that of the BD field, and would thus have a minimum Lorentz force on the scattered electrons. This will produce only a small perturbation to the dose distribution received by the patient. This effect has been studied, and is described by Bielajew, Med. Phys, vol 20, no. 4, pp 1171-1179 (1993).
The above-described patent to Green however, suggests an embodiment where the B0 field of the MRI and the direction of the x-ray beam are parallel. A defining feature of disclosure Green, however, is that it uses a standard linear accelerator configuration where the accelerating waveguide is mounted on a C-arm gantry, and the accelerating waveguide is parallel to the floor, and rotates about an axis that is also parallel to the floor. Further, the layout of the MRI and accelerating waveguide described in Green is such that the linear accelerator uses a 270° betiding magnet to direct the photon beam toward the MRI. While contemplated, it would be understood by the skilled worker that such an embodiment is, however, highly unpractical. For example, an MRI that produces images of human subjects with a field of view that is large enough and has sufficient contrast to be useful in image guided radiotherapy is far larger than those that can fit directly under a standard linear accelerator as is suggested by Green. This is simply because the size of the MRI magnet is strongly related to the desired field of view size, and contrast is directly related to magnetic field strength. In other words, systems built according to the proportions shown in the Figures of Green would simply not be capable of producing images useful for guiding radiotherapy because it could not support magnets required to do so.
A further difficulty with Green is that it clearly relies on a low magnetic field strength to reduce magnetic interference between the MRI and linear accelerator. Further, Green suggests methods whereby NMR spectral techniques are used to visualize a radiation beam. However, those skilled in the art, knowing that the magnetic field strength is a limiting factor when producing high contrast imaging, would immediately recognize that one cannot rely on a low magnetic field to produce MRI images that are in arty way suitable for guiding radiotherapy. Further, it is well known that NMR spectroscopy functions well only at high magnetic field strengths.
It is therefore an object of the invention to at least mitigate the disadvantages encountered when integrating a linear accelerator and an MRI for image guided radiotherapy.